This invention relates generally to medical imaging systems, and more particularly to computed tomography (CT). Although the application subject matter finds particular use in x-ray systems, the invention may also find use in connection with other imaging modalities.
Modern CT scanners typically employ thousands of x-ray detectors to convert x-ray energy to electrical signals. A typical detector may include an array of scintillators attached to an array of semiconductor photodiodes which detect light or other ionizing radiation on their front surface. Some implementations have configurable detectors wherein signal currents from multiple individual photodiodes can be combined for further processing in a single amplifier channel. This arrangement permits the detection area for an individual pixel to be varied, using externally controlled electrical switches (field effect transistors, or FETs). The bond pads, for electrical connection to the FETs, are typically located at one or both ends of the photodiode, and the entire pixel array must be channeled from the center of the array toward one or both edges near the FETs.
As the number of elements in arrays increases, the density of the traces and bond pads increases to an unattractively high level near the edge of the photodiode array. This places some physical limits on the number and size of traces and bond pads that can be made using top surface contacts. With available wire bonding and silicon processing technology, no more than 40-50 slices of 0.625 mm pixels (measured at iso-center of a CT gantry) can be achieved.
Herein described are methods and apparatus which at least partially overcome the above-referenced problems.